MR imaging of internal body tissues may be used for numerous medical procedures, including diagnosis and surgery. In general terms, MR imaging starts by placing a subject in a relatively uniform, static magnetic field. The static magnetic field causes hydrogen nuclei spins to align and precess about the general direction of the magnetic field. Radio frequency (RF) magnetic field pulses are then superimposed on the static magnetic field to cause some of the aligned spins to alternate between a temporary high-energy non-aligned state and the aligned state, thereby inducing an RF response signal, called the MR echo or MR response signal. It is known that different tissues in the subject produce different MR response signals, and this property can be used to create contrast in an MR image. An RF receiver detects the duration, strength, and source location of the MR response signals, and such data are then processed to generate tomographic or three-dimensional images.
MR imaging can also be used effectively during a medical procedure to assist in locating and guiding medical instruments. For example, a medical procedure can be performed on a patient using medical instruments while the patient is in an MRI machine. The medical instruments may be for insertion into a patient or they may be used externally but still have a therapeutic or diagnostic effect. For instance, the medical instrument can be an ultrasonic device, which is disposed outside a patient's body and focuses ultrasonic energy to ablate or necrose tissue or other material on or within the patient's body. The MRI machine preferably produces images at a high rate so that the location of the instrument (or the focus of its effects) relative to the patient may be monitored in real-time (or substantially in real-time). The MRI machine can be used for both imaging the targeted body tissue and locating the instrument, such that the tissue image and the overlaid instrument image can help track an absolute location of the instrument as well as its location relative to the patient's body tissue.
MR imaging can further provide a non-invasive means of quantitatively monitoring in vivo temperatures. This is particularly useful in the above-mentioned MR-guided focused ultrasound (MRgFUS) treatment or other MR-guided thermal therapy where temperature of a treatment area should be continuously monitored in order to assess the progress of treatment and correct for local differences in heat conduction and energy absorption. The monitoring (e.g., measurement and/or mapping) of temperature with MR imaging is generally referred to as MR thermometry or MR thermal imaging.
Among the various methods available for MR thermometry, the proton-resonance frequency (PRF) shift method is often preferred due to its excellent linearity with respect to temperature change, near-independence from tissue type, and good sensitivity. The PRF shift method is based on the phenomenon that the MR resonance frequency of protons in water molecules changes linearly with temperature. Since the frequency change is small, only −0.01 ppm/° C. for bulk water and approximately −0.0096 to −0.013 ppm/° C. in tissue, the PRF shift is typically detected with a phase-sensitive imaging method in which the imaging is performed twice: first to acquire a baseline phase image prior to a temperature change and then to acquire a second phase image after the temperature change, thereby capturing a small phase change that is proportional to the change in temperature.
A phase image, for example, may be computed from MR image data, and a temperature-difference map relative to the baseline image may be obtained by (i) subtracting, on a pixel-by-pixel basis, the phase image corresponding to the baseline from the phase image corresponding to a subsequently obtained MR image, and (ii) converting phase differences into temperature differences based on the PRF temperature dependence.
Unfortunately, changes in phase images do not arise uniquely from temperature changes. Various non-temperature-related factors, such as changes in a local magnetic field due to nearby moving metal, magnetic susceptibility changes in a patient's body due to breathing or movement, and magnet or shim drifts can all lead to confounding phase shifts that may render a phase-sensitive temperature measurement invalid. The changes in magnetic field associated with magnet drift and patient motion are often severe enough to render temperature measurements made using the above-mentioned phase-sensitive approach useless.
Spurious phase shifts can be quite significant when temperature changes are monitored over a long time period, such as during a lengthy treatment procedure. As the elapsed time between the initial baseline phase image and the actual temperature measurement increases, concurrent (and non-temperature-related) changes in magnetic field are more likely to occur, impairing the accuracy of temperature measurement. For example, in existing MR-guided thermal treatment procedures, it is often assumed that the main magnetic field and gradient field are sufficiently stable during the treatment and that the pre-treatment temperature of a target area is known, such that any phase shift is assumed to be due exclusively to change in temperature. These assumptions might be valid in certain MRgFUS procedures where tissues of interest lie deep within a patient's body or where heating periods are short (e.g., less than a minute per heating period, followed by a cooling-down period allowing the tissues to return to body temperature). However, the above-mentioned assumptions do not hold up when heating periods are relatively long or with slower heating methods (e.g., radiofrequency and laser heating). Nor are these assumptions valid when an initial tissue temperature is unknown, such as when the treatment area is close to skin surface or is actively cooled.
Also, in some applications of MR thermometry, it may be critical or desirable to measure absolute temperature(s) instead of a simple change in temperature. For example, in a prostate treatment with MRgFUS, an absolute temperature of the treatment area may be required in order to accurately calculate a thermal dose. However, if the patient's rectum is actively cooled for safety reasons, there will be a gradual temperature change between the cooled rectal wall and the inner tissue of the prostate. In general, this temperature profile cannot be estimated with sufficient accuracy due to variability of tissue properties among patients (e.g., differences in perfusion), so an absolute temperature measurement becomes necessary. Another example relates to an MRgFUS treatment of soft tissue tumors, where a significant amount of time may be spent waiting for a heated tissue to cool down after each delivery of ultrasonic energy. Although long cooling periods may not be necessary for safety or efficacy reasons, they are still deemed necessary because the temperature measurement during the next energy delivery (sonication) relies on the assumption that the heated tissue has returned to body temperature. If an absolute temperature of the heated tissue were measured, subsequent sonications could start sooner and thus the overall treatment time could be significantly shortened.
In view of the foregoing, it may be understood that there are significant problems and shortcomings associated with current PRF techniques.